This invention relates to reciprocating pump arrangements for pumping fluids such as blood in a controlled manner. More specifically, this invention is directed to a reciprocating pump capable of providing optimal assistance for ventricular or cardiac support such as that for an ailing ventricle. It produces effective pumping action under minimum shear conditions.
Heretofore a number of pump designs have been proposed for pumping fluids such as blood. Such pumps must provide leak-free operation and must avoid contamination of the fluid by the pump components and the external environment. In addition, such pumps must effectively pump the fluid at a suitable rate without applying excessive Reynolds shear stress to the fluid. Damage due to excessive shear is particularly a problem when pumping fluids such as blood or blood products.
It is well known to those skilled in the art that lysis or cell destruction may result from application of shear stress to cell membranes. Red blood cells are particularly susceptible to shear stress damage as their cell membranes do not include a reinforcing cytoskeleton to maintain cell shape. Lysis of white blood cells and platelets also occurs upon application of high shear stress. Lysis of red blood cells can result in release of cell contents which trigger subsequent platelet aggregation. Sublytic shear stress leads to cellular alterations and direct activation and aggregation of platelets and white blood cells.
Several types of surgically implantable pumps have been developed in an effort to provide a mechanical device for augmenting or replacing the blood pumping action of damaged or diseased hearts. Some of these pumps are designed to support single ventricular function. Such pumps usually support the left ventricle, which pumps blood to the entire body except the lungs, since it becomes diseased far more commonly than the right ventricle, which pumps blood only to the lungs. Other devices have been tested and used for providing biventricular function.
Depending on the needs of a particular patient and the design of a pump, pumping units such as so-called "VADs" (ventricular assist devices) can be implanted to assist a functioning heart that does not have adequate: pumping capability. Other types of pumps, such as the so-called "Jarvik heart," can be used to completely replace a heart which has been surgically removed.
Temporary as well as permanent implantable pumps have been developed. "Permanent" in this sense refers to the remaining life of the patient; after a patient's death, any artificial pumping device is usually removed for analysis. "Temporary" implantation usually involves (1) an attempt to reduce the stress on a heart while it recovers from surgery or some other short-term problem, or (2) use of a pump as a "bridge" to forestall the death of a patient until a suitable donor heart can be found for cardiac transplantation.
The most widely tested and commonly used implantable blood pumps employ variable forms of flexible sacks (also spelled sacs) or diaphragms which are squeezed and released in a cyclical manner to cause pulsatile ejection of blood. Such pumps are discussed in books or articles such as Hogness and Antwerp 1991, DeVries et al 1984, and Farrar et al 1988, and in U.S. Pat. Nos. 4,994,078 (Jarvik 1991), 4,704,120 (Slonina 1987), 4,936,758 (Coble 1990), and 4,969,864 (Schwarzmann et al 1990). Sack or diaphragm pumps are subject to fatigue failure of compliant elements and as such are mechanically and functionally quite different from the pump which is the subject of the present invention.
An entirely different class of implantable blood pumps uses rotary pumping mechanisms. Most rotary pumps can be classified into two categories: centrifugal pumps and axial pumps. Centrifugal pumps, which include pumps marketed by Sarns (a subsidiary of the 3M Company) and Biomedicus (a subsidiary of Medtronic, Eden Prairie, Minn.), direct blood into a chamber, against a spinning interior wall (which is a smooth disk in the Medtronic pump). A flow channel is provided so that the centrifugal force exerted on the blood generates flow.
By contrast, axial pumps provide blood flow along a cylindrical axis, which is in a straight (or nearly straight) line with the direction of the inflow and outflow. Depending on the pumping mechanism used inside an axial pump, this can in some cases reduce the shearing effects of the rapid acceleration and deceleration forces generated in centrifugal pumps. However, the mechanisms used by axial pumps can inflict other types of stress and damage on blood cells.
Some types of axial rotary pumps use impeller blades mounted on a center axle, which is mounted inside a tubular conduit. As the blade assembly spins, it functions like a fan, or an outboard motor propeller. As used herein, "impeller" refers to angled vanes (also called blades) which are constrained inside a flow conduit; an impeller imparts force to a fluid that flows through the conduit which encloses the impeller. By contrast, "propeller" usually refers to non-enclosed devices, which typically are used to propel vehicles such as boats or airplanes.
Another type of axial blood pump, called the "Haemopump" (sold by Nimbus) uses a screw-type impeller with a classic screw (also called an Archimedes screw; also called a helifoil, due to its helical shape and thin cross-section). Instead of using several relatively small vanes, the Haemopump screw-type impeller contains a single elongated helix, comparable to an auger used for drilling or digging holes. In screw-type axial pumps, the screw spins at very high speed (up to about 10,000 rpm). The entire Haemopump unit is usually less than a centimeter in diameter. The pump can be passed through a peripheral artery into the aorta, through the aortic valve, and into the left ventricle. It is powered by an external motor and drive unit.
Centrifugal or axial pumps are commonly used in three situations: (1) for brief support during cardio-pulmonary operations, (2) for short-term support while awaiting recovery of the heart from surgery, or (3) as a bridge to keep a patient alive while awaiting heart transplantation. However, rotary pumps generally are not well tolerated for any prolonged period. Patients who must rely on these units for a substantial length of time often suffer from strokes, renal (kidney) failure, and other organ dysfunction. This is due to the fact that rotary devices, which must operate at relatively high speeds, may impose unacceptably high levels of turbulent and laminar shear forces on blood cells. These forces can damage or lyse (break apart) red blood cells. A low blood count (anemia) may result, and the disgorged contents of lysed blood cells (which include large quantities of hemoglobin) can cause renal failure and lead to platelet activation that can cause embolisms and stroke.
One of the most important problems in axial rotary pumps in the prior art involves the gaps that exist between the outer edges of the blades, and the walls of the flow conduit. These gaps are the site of severe turbulence and shear stresses, due to two factors. Since implantable axial pumps operate at very high speed, the outer edges of the blades move extremely fast and generate high levels of shear and turbulence. In addition, the gap between the blades and the wall is usually kept as small as possible to increase pumping efficiency and to reduce the number of cells that become entrained in the gap area. This can lead to high-speed compression of blood cells as they are caught in a narrow gap between the stationary interior wall of the conduit and the rapidly moving tips or edges of the blades.
An important factor that needs to be considered in the design and use of implantable blood pumps is "residual cardiac function," which is present in the overwhelming majority of patients who would be candidates for mechanical circulatory assistance. The patient's heart is still present and still beating, even though, in patients who need mechanical pumping assistance, its output is not adequate for the patient's needs. In many patients, residual cardiac functioning often approaches the level of adequacy required to support the body, as evidenced by the fact that the patient is still alive when implantation of an artificial pump must be considered and decided. If cardiac function drops to a level of severe inadequacy, death quickly becomes imminent, and the need for immediate intervention to avert death becomes acute.
Most conventional ventricular assist devices are designed to assume complete circulatory responsibilities for the ventricle they are "assisting." As such, there is no need, nor presumably any advantage, for the device to interact in harmony with the assisted ventricle. Typically, these devices utilize a "fill-to-empty" mode that, for the most part, results in emptying of the device in random association with native heart contraction. This type of interaction between the device and assisted ventricle ignores the fact that the overwhelming majority of patients who would be candidates for mechanical assistance have at least some significant residual cardiac function.
It is preferable to allow the natural heart, no matter how badly damaged or diseased it may be, to continue contributing to the required cardiac output whenever possible so that ventricular hemodynamics are disturbed as little as possible. This points away from the use of total cardiac replacements and suggests the use of "assist" devices whenever possible. However, the use of assist devices also poses a very difficult problem: in patients suffering from severe heart disease, temporary or intermittent crises often require artificial pumps to provide "bridging" support which is sufficient to entirely replace ventricular pumping capacity for limited periods of time, such as in the hours or days following a heart attack or cardiac arrest, or during periods of severe tachycardia or fibrillation.
Accordingly, an important goal during development of the described method of pump implantation and use and of the surgically implantable reciprocating pump was to design a method and a device which could cover a wide spectrum of requirements by providing two different and distinct functions. First, an ideal cardiac pumping device should be able to provide "total" or "complete" pumping support which can keep the patient alive for brief or even prolonged periods, if the patient's heart suffers from a period of total failure or severe inadequacy. Second, in addition to being able to provide total pumping support for the body during brief periods, the pump should also be able to provide a limited "assist" function. It should be able to interact with a beating heart in a cooperative manner, with minimal disruption of the blood flow generated by the natural heartbeat. If a ventricle is still functional and able to contribute to cardiac output, as is the case in the overwhelming majority of clinical applications, then the pump will assist or augment the residual cardiac output. This allows it to take advantage of the natural, non-hemolytic pumping action of the heart to the fullest extent possible; it minimizes red blood cell lysis, it reduces mechanical stress on the pump, and it allows longer pump life and longer battery life.
Several types of surgically implantable blood pumps containing a piston-like member have been developed to provide a mechanical device for augmenting or even totally replacing the blood pumping action of a damaged or diseased mammalian heart.
U.S. Pat. No. 3,842,440 to Karlson discloses an implantable linear motor prosthetic heart and control system containing a pump having a piston-like member which is reciprocal within a magnetic field. The piston-like member includes a compressible chamber in the prosthetic heart which communicates with the vein or aorta.
U.S. Pat. Nos. 3,911,897 and 3,911,898 to Leachman, Jr. disclose heart assist devices controlled in the normal mode of operation to copulsate and counterpulsate with the heart, respectively, and produce a blood flow waveform corresponding to the blood flow waveform of the heart being assisted. The heart assist device is a pump connected serially between the discharge of a heart ventricle and the vascular system. The pump may be connected to the aorta between the left ventricle discharge immediately adjacent the aortic valve and a ligation in the aorta a short distance from the discharge. This pump has coaxially aligned cylindrical inlet and discharge pumping chambers of the same diameter and a reciprocating piston in one chamber fixedly connected with a reciprocating piston of the other chamber. The piston pump further includes a passageway leading between the inlet and discharge chambers and a check valve in the passageway preventing flow from the discharge chamber into the inlet chamber. There is no flow through the movable element of the piston.
U.S. Pat. No. 4,102,610 to Taboada et al. discloses a magnetically operated constant volume reciprocating pump which can be used as a surgically implantable heart pump or assist. The reciprocating member is a piston carrying a tilting-disk type check valve positioned in a cylinder. While a tilting disk valve results in less turbulence and applied shear to surrounding fluid than a squeezed flexible sack or rotating impeller, the shear applied may still be sufficiently excessive so as to cause damage to red blood cells.
U.S. Pat. Nos. 4,210,409 and 4,375,941 to Child disclose a pump used to assist pumping action of the heart having a piston movable in a cylindrical casing in response to magnetic forces. A tilting-disk type check valve carried by the piston provides for flow of fluid into the cylindrical casing and restricts reverse flow. A plurality of longitudinal vanes integral with the inner wall of the cylindrical casing allow for limited reverse movement of blood around the piston which may result in compression and additional shearing of red blood cells. A second fixed valve is present in the inlet of the valve to prevent reversal of flow during piston reversal.
U.S. Pat. No. 4,965,864 to Roth discloses a linear motor using multiple coils and a reciprocating element containing permanent magnets which is driven by microprocessor-controlled power semiconductors. A plurality of permanent magnets is mounted on the reciprocating member. This design does not provide for self-synchronization of the linear motor in the event the stroke of the linear motor is greater than twice the pole pitch on the reciprocating element. During start-up of the motor, or if magnetic coupling is lost, the reciprocating element may slip from its synchronous position by any multiple of two times the pole pitch. As a result, a sensing arrangement must be included in the design to detect the position of the piston so that the controller will not drive it into one end of the closed cylinder. In addition, this design having equal pole pitch and slot pitch results in a "jumpy" motion of the reciprocating element along its stroke.
In addition to the piston position sensing arrangement discussed above, the Roth design may also include a temperature sensor and a pressure sensor as well as control circuitry responsive to the sensors to produce the intended piston motion. For applications such as implantable blood pumps where replacement of failed or malfunctioning sensors requires open heart surgery, it is unacceptable to have a linear motor drive and controller that relies on any such sensors. In addition, the Roth controller circuit uses only NPN transistors thereby restricting current flow to the motor windings to one direction only.
U.S. Pat. No. 4,541,787 to Delong describes a pump configuration wherein a piston containing a permanent magnet is driven in a reciprocating fashion along the length of a cylinder by energizing a sequence of coils positioned around the outside of the cylinder. However, the coil and control system configurations disclosed only allow current to flow through one individual winding at a time. This does not make effective use of the magnetic flux produced by each pole of the magnet in the piston. To maximize force applied to the piston in a given direction, current must flow in one direction in the coils surrounding the vicinity of the north pole of the permanent magnet while current flows in the opposite direction in the coils surrounding the vicinity of the south pole of the permanent magnet. Further, during starting of the pump disclosed by Delong, if the magnetic piston is not in the vicinity of the first coil energized, the sequence of coils that are subsequently energized will ultimately approach and repel the magnetic piston toward one end of the closed cylinder. Consequently, the piston must be driven into the end of the closed cylinder before the magnetic poles created by the external coils can become coupled with the poles of the magnetic piston in attraction.
U.S. Pat. No. 4,610,658 to Buchwald et al. discloses an implantable fluid displacement peritoneovenous shunt system. The system comprises a magnetically driven pump having a spool piston fitted with a disc flap valve.
U.S. Pat. No. 5,089,017 to Young et al. discloses a drive system for artificial hearts and left ventricular assist devices comprising one or more implantable pumps driven by external electromagnets.